The invention relates generally to implantable blood pumps, and particularly to a fully implantable single chamber blood pump apparatus.
Serious heart failure, or the inability of a person's heart to pump sufficient blood for their body's needs, is the cause of very poor quality of life, huge medical treatment costs, and death in hundreds of thousands of patients yearly. Numerous pharmacologic, biologic, and device interventions have been devised to deal with this disease, many of them patented, but despite these efforts, heart failure remains a major public health problem.
The measure of heart failure is an abnormally low cardiac output or cardiac index. Cardiac output (CO) is measured in liters of blood flow per minute (l/min) and cardiac index (CI) is CO divided by the patient's body surface area (BSA). Normally CI at rest or during light activity is between 3.0 and 3.5 and CO is between 5.6 to 6.5 liters per minute for men and proportionally less for women based upon less body surface area Severe heart failure exists when the CI is between 1.5 and 2.0. For an average man in heart failure with 1.87 meters squared BSA, a cardiac index of 1.75 and a heart rate of 80 beats per minute (BPM), the cardiac output will be 3.27 l/min and an average of 41 ml of blood will be ejected from the heart with each heartbeat. This average stroke volume contrasts with an average normal stroke volume of 76 ml, which would occur in an average normal man with a CI of 3.25 and heart rate of 80 BPM.
The main pumping chamber of the heart or left ventricle (LV), has an inlet (mitral) valve and an outlet (aortic) valve. During left ventricluar contraction, the inlet valve closes as blood is pushed through the aortic valve and into the aorta or main artery of the body. Resting (diastolic) LV pressure may be between 2 and 20 mm Hg pressure (preload) and will be in the higher end of this range during failure. During active LV contraction (systole), the LV must eject the blood against aortic pressure, which is typically between 70 and 140 mm Hg of pressure (afterload). It is well known that, if in failure the afterload is reduced, the stroke volume will naturally increase and this increase is one reason that afterload-reducing drugs such as ACE-inhibitors have helped heart failure patients.
A common method of providing mechanical circulatory assist is the use of counterpulsation devices such as intraaortic balloon pumps (IABPs). IABPs provide an afterload-reducing type of assist and are typically employed for acute use (i.e. for hours to days). As described in U.S. Pat. Nos. 4,733,652 and 3,692,018 to Kantrowitz et al. and Goetz et al., respectively, the main benefit of such devices stems from unloading the left ventricle during systole and providing increased diastolic pressure for reperfusing the coronary and other arteries during diastole. Patients needing this type of treatment suffer from cardiogenic shock, chronic angina, or need perioperative circulatory support (Nanas et al. 1988, Kormos 1987). The nature of IABP design restricts itself to acute use only, since the bulky balloon drive unit remains outside the patient's body necessitating confinement to a hospital bed.
Pouch-type auxiliary ventricles that have mechanical or pneumatic means for the pumping the contained blood are disclosed in U.S. Pat. Nos. 3,553,736 and 4,034,742 to Kantrowitz et al. and Thoma, respectively. Many of these have a single access port that serves as both the inlet and the outlet for bloodflow. These designs have the disadvantage of relative flow stagnation which increases the risk of clot formation and thromboembolism. Others have both an inlet and outlet port and can be connected in parallel with the aorta. These designs may have valves to attempt to maximize their pumping effectiveness (U.S. Pat. Nos. 4,195,623 and 4,245,622 to Zeffet al. and Hutchins, respectively.)
A "dynamic aortic patch" is disclosed in U.S. Pat. Nos. 4,630,596 and 4,051,840, both to Kantrowitz et al., which is permanently attached to the aorta and is designed to provide counterpulsation assistance. This device is intended for chronic use and requires opening the patient's thorax for installation. Like the IABP, the drive unit remains outside the patient's body and inflation of the patch is accomplished pneumatically through a percutaneous access port. Unlike the IABP, the dynamic aortic patch can produce volumetric assistance greater than 40 ml. Two risks of this system are the risk of chronic infection due to the permanent percutaneous port and the extensiveness of the implant surgery. Physically, the patch is oblong in shape and consists of a flexible balloon on the blood side of a chamber that has a rigid back through which a pneumatic line (hose) passes to effect balloon inflation and deflation. Along the perimeter of the rigid back is a flange that provides an edge for suturing the patch into the aortic wall. The hose penetrates the skin surface percutaneously through a specially designed skin port. Inflation and deflation of the balloon is accomplished by an external air pump that is connected to the intraaortic patch during operation. When and if the balloon is not being pulse driven, the aorta is open to bloodflow allowing the pump to be fail-safe. In the standby mode, the balloon interior is then at atmospheric pressure which is lower than aortic blood pressure causing the pumping chamber to collapse.
The addition of compliance to the arterial system is described in U.S. Pat. No. 4,938,766 to Jarvik. Hardening of the arteries lowers vascular compliance and can increase the afterload presented to the heart. Consequently, the addition of a compliance chamber can help to somewhat reverse the effects of arterial hardening and hence decreases the heart's workload. From the disclosure, such devices are generally used to assist the left ventricle. Several configurations of compliance chambers are disclosed and various methods of implantation are also taught. The devices can be categorized as single-port chambers, two-port flow-though chambers, and spring-loaded mechanical clips that are attached to the aorta. For designs having a flow-through configuration, a valve may be included in the inlet side of the chamber. This can be for preventing backflow and preferentially can direct the outflow of blood from the compliance chamber towards more desired locations.
Direct pumping during heart diastole is typically performed by what are referred to as ventricular assist devices (VADs). VADs that have a flow-through configuration and which convert electric energy directly to mechanical energy are the most pertinent prior art to the invention described herein. U.S. Pat. No. 4,091,471 to Richter describes an apparatus that mechanically compresses a toroidal flow conduit by squeezing the inner radius and pushing it outward while preventing the outer radius from expanding. This is accomplished through pressurization of a sealed central portion, in the center of the toroid. U.S. Pat. No. 4,250,872 to Tamari describes a flow-through pumping chamber which is squeezed by a pressurizing fluid. The Tamari device relies mainly on thickness variations of the pumping chamber wall to control compression of the pumping chamber. U.S. Pat. No. 5,089,016 to Millner et al. has a flow-through toroidal design which employs a hydraulic pumping fluid to compress the pumping toroidal chamber. The Millner device can have valves at both the inlet and the outlet of the pumping chamber. The pumping chamber itself is squeezed from all directions circumferentially to accomplish blood pumping. However, to minimize wall stress in the pumping chamber, the chamber can preferably be squeezed in one direction by stiffening the opposite side of the chamber wall.
Articles published by Frazier et al. (Circulation, 89:2908-2914, 1994) and McCarthy et al. (Ann Thoracic Surg, 59:S46-S51, 1995) describe a blood pump which has a diaphragm-driven circular chamber that is implanted in the upper left abdominal wall and is capable of an 83 ml stroke volume. The pumping chamber receives blood from a conduit that pierces the apex of the LV. The pumping diaphragm may be driven pneumatically or by an electric motor driving a single rotation roller-cam mechanism. In both cases, the pumping chamber is circular and drive lines pierce the skin. Adequate filling of the chamber is possible because the nonblood side of the drive membrane is vented to the atmosphere via the skin port.
The blood pump disclosed in U.S. Pat. No. 5,569,156 to Mussivand has the nonblood side of the drive membrane contacting hydraulic fluid that, during filling, must be actively pumped to a separate volume displacement chamber (VDC). The blood pump also has inlet and outlet ports that are perpendicular to the blood pumps drive membrane.
The blood pump disclosed in an article by Ramasamy et al. (ASAIO Transactions, 35:402-404, 1989) illustrates a separate gas filled compliance chamber placed in the pleural space that communicates by means of gas tight tubing to the nonblood contacting side of the blood chambers pumping membrane.
For ease of filling, it is necessary for the non-blood-contacting side of the diaphragm to be at or near atmospheric pressure to permit easy blood inflow. Artificial VAD pumping chambers, together with their associated electronics and drive mechanisms, are not yet sufficiently compact to be fully implanted. Instead, various leads have to penetrate the skin and connect the pumping chamber with the external drive mechanism. To permit easy filling of the pumping chamber with blood, a low opposing pressure is needed. The preceding articles describe three means for accomplishing this low pressure; venting through the skin to atmospheric pressure, venting to a separate gas filled compliance chamber in the pleural space, and using an intermediate hydraulic fluid connected to an implanted volume displacement chamber or VDC.
The pumping chamber described in Frazier et al. has a bloodflow path which is parallel to the LV since the chamber receives blood from the LV apex and pumps the blood into the aorta beyond with a flow path in parallel with the LV.
Accordingly, there is a need for a blood pump which is small enough that, together with its associated electronics, can be totally implanted to avoid the infection risk associated with percutaneous leads. The blood pump also should not require a second chamber for compliance, whether in the form of a chamber which is a part of the pump or a separately located chamber connected to the pump with gas tight tubing. Moreover, in contrast to the parallel connection pathway, it can be preferable to have a blood pump which receives blood from the aortic root at a low filling pressure and return the blood to the ascending aorta by driving the pump and blood to a higher pressure than exists in the distal aorta. Such a connection configuration would be referred to as "in-series" with the left ventricle.